Miniature radiation source with flexible probe and laser driven thermionic emitter

ABSTRACT

This invention is directed to a radiation source comprising a power supply, a flexible fiber optic cable assembly, a light source, and a target assembly. The power supply includes a first terminal and a second terminal, and elements for establishing an output voltage between the first terminal and the second terminal. The flexible fiber optical cable assembly has an originating end and a terminating end, and includes a fiber optical element extending from the originating end to the terminating end. The cable is adapted for transmitting light incident on the originating end to the terminating end. The light source includes elements for generating a beam of light at and directed to the originating end of the fiber optical cable assembly. The target assembly is affixed to the terminating end of the fiber optical cable assembly and is electrically coupled to the power supply by way of the first terminal and the second terminal. The target assembly includes elements for emitting radiation in a predetermined spectral range, in response to light transmitted to the terminating end.

BACKGROUND OF THE INVENTION

[0001] The present invention relates to a highly miniaturized, lowpower, programmable radiation source for use in delivering predefineddoses of radiation to a predefined region and more particularly to aminiaturized radiation source mounted in a flexible probe.

[0002] In the field of medicine, radiation is used for diagnostic,therapeutic and palliative treatment of patients. The conventionalmedical radiation sources used for these treatments include large fixedposition machines as well as small, transportable radiation generatingprobes. The current state of the art treatment systems utilize computersto generate complex treatment plans.

[0003] These systems apply doses of radiation that are known to inhibitthe growth of new tissue because the radiation affects dividing cellsmore than the mature cells found in non-growing tissue. Thus, the tissuein the site of an excised tumor can be treated to prevent the regrowthof cancerous tissue and the recurrence of cancer. Alternatively,radiation can be applied to other areas of the body to inhibit tissuegrowth, for example the growth of new blood vessels inside the eye thatcan cause macular degeneration.

[0004] Conventional radiation treatments systems, such as the LINAC usedfor medical treatment, utilize a high power remote radiation source anddirect a beam of radiation at a target area, such as tumor inside thebody of a patient. This type of treatment is referred to as teletherapybecause the radiation source is located a predefined distance,approximately one meter, from the target. This treatment suffers fromthe disadvantage that tissue disposed between the radiation source andthe target is exposed to radiation.

[0005] An alternative treatment system utilizing a point source ofradiation is disclosed in U.S. Pat. No. 5,153,900 issued to Nomikos etal., U.S. Pat. No. 5,369,679 to Sliski et al., and U.S. Pat. No.5,422,926 to Smith et al., all owned by the assignee of the presentapplication, all of which are hereby incorporated by reference. Thissystem includes a miniaturized, insertable probe capable of producinglow power radiation in predefined dose geometries disposed about apredetermined location. This type of treatment is referred to asbrachytherapy because the source is located close to or in some caseswithin the area receiving treatment. One advantage of brachytherapy isthat the radiation is applied primarily to treat a predefined tissuevolume, without significantly affecting the tissue adjacent to thetreated volume.

[0006] Typical radiation therapy treatment involves positioning theinsertable probe into or adjacent to the tumor or the site where thetumor or a portion of the tumor was removed to treat the tissue adjacentthe site with a “local boost” of radiation. In order to facilitatecontrolled treatment of the site, it is desirable to support the tissueportions to be treated at a predefined distances from the radiationsource. Alternatively, where the treatment involves the treatment ofsurface tissue or the surface of an organ, it is desirable to controlthe shape of the surface as well as the shape of the radiation fieldapplied to the surface.

[0007] The treatment can involve the application of radiation, eithercontinuously or intermittently, over an extended period of time.Therefore, it is desirable that the insertable probe be adjustablysupported in a compliant manner to accurately position the radiationsource with respect to the treated site and accommodate normal minormovements of the patient, such as movements associated with breathing.

[0008] In many x-ray therapeutic procedures, x-ray probes of the typegenerally disclosed in U.S. Pat. No. 5,153,900 incorporate a relativelyrigid tube enclosing an electron beam directed to an x-ray emittingtarget at its distal end. For example, in treatment of brain tumors, anx-ray probe having a rigid tube is used with a stereotactic frameaffixed to the patient's skull, where the tube is advanced into a biopsyhole to the tumor location, as disclosed in U.S. Pat. No. 5,369,679. Therigidity of the tube is useful in ensuring that the x-ray emittingtarget is properly located. In other cases, it is desirable to have aflexible tube leading to the x-ray emitting target, for example, whereit is desirable to pass the probe up the urethra to the bladder, fortreatment of the bladder. Such a flexible probe is disclosed in U.S.Pat. No. 5,248,658.

[0009] However, it has been difficult to effectively treat tissue usingthe flexible probe of the latter patent.

[0010] Accordingly, it is an object of the present invention to providean improved system for delivering radiation to a localized area.

[0011] It is a further object of the present invention to provide animproved highly miniaturized radiation source with a flexible probe.

SUMMARY OF THE INVENTION

[0012] The present invention is directed to a miniaturized radiationsource at the end of a flexible probe or catheter. The flexible catheterextends along a probe axis between a proximal end and a distal end ofthe catheter. The radiation source, at the distal end of the catheter,includes a substantially rigid housing disposed about a substantiallyevacuated interior region extending along a beam axis from an electronsource at an input end of the housing to a radiation transmissive windowat an output end of the housing. The housing also includes a channelelectron multiplier adapted for receiving electrons from the electronsource and for producing free electrons at an output end of the channelelectron multiplier and an electron accelerator adapted for establishinga potential difference in the interior region of the housing whereby thefree electrons produced at the output end of the channel electronmultiplier are accelerated toward a target at or near the window. Thetarget produces x-radiation in response to incident accelerated freeelectrons.

[0013] Preferably, the electron accelerator includes a surface disposedabout the beam axis between the electron source and the target on aceramic and preferably monolithic, substrate. In one embodiment, thesurface bears a semiconductor coating. The surface may be substantiallyconical in shape wherein the distance from the beam axis increases as afunction of the distance from the electron source. The electron sourcecan be a photocathode illuminated by laser energy, a field emitter or athermionic emitter. The target and outer surface of the probe ispreferably maintained at ground potential to reduce the risk of shock.

BRIEF DESCRIPTION OF THE DRAWINGS

[0014] The foregoing and other objects of this invention, the variousfeatures thereof, as well as the invention itself, may be more fullyunderstood from the following description, when read together with theaccompanying drawings in which:

[0015]FIGS. 1A and 1B are a diagrammatic perspective view and adiagrammatic detail view, respectively, of a low power radiation sourceembodying the present invention;

[0016]FIGS. 2A and 2B are a perspective view and a sectional view,respectively, of an alternate form embodying the present invention;

[0017]FIG. 3 is a diagrammatic representation of a sheath adapted foruse with the apparatus of FIG. 1;

[0018]FIG. 4 is a schematic block diagram of the embodiment of FIG. 1;

[0019]FIG. 5 is a diagrammatic view of a low power radiation treatmentsystem having a flexible probe embodying the present invention; and

[0020]FIG. 6 is a diagrammatic view of a low power radiation sourceembodying the present invention.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0021] The present invention is directed to a miniature, low powerradiation producing probe which can be used for diagnostic, therapeuticand palliative treatment of patients. The radiation source in accordancewith the present invention can made smaller than conventional radiationsources. In addition, the radiation source can be disposed at the distalend of the tip of a flexible (or rigid) tube or catheter which can beinserted into the body. In one embodiment, only a single high voltagewire is necessary for operation. The target and the outer surface of theprobe are set at the ground potential to reduce the shock hazard of thedevice.

[0022]FIG. 1A shows an x-ray apparatus 10 embodying the presentinvention. Apparatus 10 includes a housing 12 and an elongatedcylindrical probe 14 extending from housing 12 along a reference axis 16to an x-ray source assembly 19. Preferably, the probe 14 is flexible, asdescribed below, but in some cases may be rigid. The housing 12 enclosesa high voltage power supply 12A, a battery 12B and a control system 12C.The x-ray source assembly 19 has an electron source (cathode) 22 locatedin the distal end of the probe 14. Electron source 22 is located inclose proximity to a channel electron multiplier (CEM) 23 which receiveselectrons from the electron source 22. An accelerator 24 is positionedbetween the CEM 23 and an x-ray emitting (in response to incidentaccelerated electrons) target 26. The target underlies on x-raytransmissive window 27. Probe 14 is integral with the housing 12 andextends toward the target 26. In various embodiments, the x-ray emittingtip may be selectively shielded to control the spatial distribution ofx-rays. In addition, the accelerator 24 may be magnetically shielded toprevent external magnetic fields from deflecting the beam away fromdesigned impact points on the target.

[0023]FIG. 1B shows an x-ray source assembly 19′ for generating x-raysembodying the present invention. That source 19′ is adapted forplacement at the end of a cylindrical element (flexible or rigid). In analternate form, shown in FIG. 2A and 2B, an x-ray source 19″ ispositioned within a compact housing 12, the latter device is suitable toapplying x-radiation body surface tissue.

[0024] In the various forms of x-ray source assembly 19, the electronbeam generator 22 may include a thermionic emitter (driven by a lowvoltage power source or laser) or a photocathode (irradiated by an LEDor laser source) or a field emitter. A single high voltage power supply12A can be used to power the electron source (thermionic emitter) 22,the CEM 23 and accelerator 24. The accelerator 24 establishes anacceleration potential difference between the CEM 23 and the target 26which is at ground potential. The beam generation and accelerationcomponents can be adapted to establish a thin (e.g. 1 mm or less indiameter) electron beam within the assembly 19 along a nominallystraight axis 16.

[0025] Preferably, the CEM 23 is constructed as is well known and theelectron multiplication value is predetermined as function of theintended use of the radiation source. Preferably, a high voltage of 1 Kvolt is connected to input end of the CEM.

[0026] Preferably, the accelerator is constructed from a monolithicceramic material and includes an interior channel formed in the shape ofthe surface of a cone, although other shapes may be used, for exampleparabolic. The accelerator is disposed between the CEM 23 and the target26 along the axis of the electron beam trajectory whereby the distanceof the surface from the beam increases as a function of the distancefrom the CEM 23. Preferably, the surface includes a semiconductivecoating 24A which ensures that the voltage gradient in the acceleratoris smooth and linear and helps to prevent breakdown which occurs whenthe electrons hit the walls of the accelerator.

[0027] In one form of the invention, the outer cylindrical portion ofthe x-ray source assembly 19 is a hollow evacuated cylinder made of amolybdenum-rhenium, (Mo-Re), molybdenum (Mo) or mu-metal body with aninterior diameter of 2 mm, and an exterior diameter of 3 mm. Preferably,beryllium (Be) cap and having a distance from the electron source to thetarget is less than 2 mm. The target assembly 26 includes an emissionelement consisting of a small beryllium (Be) target element 26A which iscoated on the side exposed to the incident electron beam with a thinfilm or layer 26B of a high-Z element, such as tungsten (W), uranium (U)or gold (Au). By way of example, with electrons accelerated to 30 keV-,a 2.2 micron thick tungsten film absorbs substantially all the incidentelectrons, while transmitting approximately 95% of any 30 keV-, 88% ofany 20 keV-, and 83% of any 10 keV- x-rays generated in that layer. Inthe preferred embodiment, the beryllium target element 26A is 0.5 mmthick with the result that 95% of the x-rays generated in directionsnormal to and toward the substrate 26A, and having passed through thetungsten target, are then transmitted through the beryllium substrateand outward at the distal end of assembly 19. While the target element26A shown in FIG. 3B is in the form of a hemispherical layer, othershaped elements may be used, such as those having disk-like or conicalshapes.

[0028] In some forms of the target, the window element 26A may include amultiple layer film 26B, where the differing layers may have differentemission characteristics. By way of example, the first layer may have anemission (vs. energy) peak at a relatively low energy, and the second(underlying) layer may have an emission (vs. energy) peak at arelatively high energy. With this form of the invention, a low energyelectron beam may be used to generate x-rays in the first layer (toachieve a first radiation characteristic) and high energy electrons maybe used to penetrate through to the underlying layer (to achieve asecond radiation characteristic). As an example, a 0.5 mm wide electronbeam is emitted at the cathode and accelerated to 30 keV- through theanode, with 0.1 eV transverse electron energies, and arrives at thetarget assembly 26 downstream from the anode, with a beam diameter ofless than 1 mm at the target assembly 26. X-rays are generated in thetarget assembly 26 in accordance with preselected beam voltage, current,and target element 26B composition. The x-rays thus generated passthrough the beryllium target substrate 26A with minimized loss inenergy. As an alternative to beryllium, the target substrate 26A may bemade of carbon or other suitable material which permits x-rays to passwith a minimum loss of energy. An optimal material for target substrate26A is carbon in its diamond form, since that material is an excellentheat conductor. Using these parameters, the resultant x-rays havesufficient energy to penetrate into soft tissues to a depth of acentimeter or more, the exact depth dependent upon the x-ray energydistribution.

[0029] The apparatus of FIGS. 2A and 2B is particularly adapted for fullimplantation into a patient, where the housing 12 a biocompatible outersurface and encloses both a high voltage power supply circuit 12A forestablishing a drive voltage for the beam generator 22, and anassociated battery 12B for driving that circuit 12A. In this case, anassociated controller 12C establishes control of the output voltage ofthe high power supply circuit 12A, in the manner described below.

[0030] The apparatus of FIGS. 1A and 1B may be used in a manner whereonly the probe 14 and x-ray source assembly 19 are inserted into apatient while the housing 12 remains outside the patient, i.e., atranscutaneous form. In the latter form, some or all of the variouselements shown within housing 12 may alternatively be remotely located.

[0031] In the transcutaneous form, the apparatus 10 may be used with anelongated closed end (or cup-shaped) sheath 34, as shown in FIG. 3,having a biocompatible outer surface, for example, fabricated of medicalgrade aliphatic polyurethane, as manufactured under the trademarkTecoflex by Thermedics, Inc., Woburn, Mass. With this configuration, theprobe 14 is first inserted into the sheath 34. The sheath 34 and probe14 are then inserted into the patient through the skin. Alternatively, aport may be inserted through the skin and attached to it, as for examplea Dermaport port manufactured by Thermedics Inc., Woburn, Mass. Theprobe 14 is then inserted into the port.

[0032] The lining of the sheath or port can be configured as an x-rayshield by introducing barium sulfate or bismuth trioxide, or other x-rayshielding materials, into the sheath. If necessary, the probe 14 andhousing 12 can be secured to the patient's body to prevent any relativemotion during the extended time of treatment. An exemplary sheath 34 isshown in FIG. 3.

[0033] In one embodiment of the apparatus as shown in FIGS. 1A and 1B,the main body of the probe 14 can be made of a magnetically shieldingmaterial such as a mu-metal. Alternatively, the probe 14 can be made ofa non-magnetic metal, preferably having relatively high values forYoung's modulus and elastic limit. Examples of such material includemolybdenum, rhenium or alloys of these materials. The outer cylindricalshell of the accelerator 24 can be made of the outer shell metal. Theinner or outer surface of probe 14 can then be coated with a highpermeability magnetic alloy such as permalloy (approximately 80% nickeland 20% iron), to provide magnetic shielding. Alternatively, a thinsleeve of mu-metal can be fitted over, or inside of that shell ofaccelerator 24. The x-ray apparatus 10 can then be used in environmentsin which there are dc and ac magnetic fields due to electrical power,the field of the earth, or other magnetized bodies nominally capable ofdeflecting the electron beam from the probe axis.

[0034] In implantable configurations, such as those of FIGS. 2A and 2B,the power supply 12A and target assembly 26 are preferably enclosed in acapsule to prevent current flow from the x-ray source to the patient.The closed housing 12 and probe 14 are, thus, encapsulated in acontinuous outer shell of appropriate shielding material such as thosementioned previously.

[0035] The high voltage power supply 12A in each of the illustratedembodiments preferably satisfies three criteria: 1) small in size; 2)high efficiency to enable the use of battery power; and 3) independentlyvariable x-ray tube voltage and current to enable the unit to beprogrammed for specific applications. A high-frequency, switch-modepower converter is used to meet these requirements. The most appropriatetopology for generating low power and high voltage is a resonant voltageconverter working in conjunction with a high voltage,Cockroft-Walton-type multiplier. Low-power dissipation, switch-modepower-supply controller-integrated circuits (IC) are currently availablefor controlling such topologies with few ancillary components.

[0036] The embodiment of FIGS. 2A and 2B can also be adapted forsuperficial usage, that is for direct placement on the skin of apatient. This form of the invention is particularly useful for x-raytreatment of skin lesions or tumors, or other dermatologicalapplications. In FIGS. 2A and 2B, elements that correspond to elementsin the embodiment of FIGS. 1A and 1B are denoted with the same referencedesignations. Apparatus 10′ generates an electron beam in a channel 40enclosed within housing 12, where that channel 40 corresponds to probe14. In the present embodiment, of FIGS. 2A and 2B, the x-ray sourceassembly 19 functions in a manner similar to that described above. Withthe configuration of FIGS. 2A and 2B, low power x-rays may be directedto a desired skin region of a patient.

[0037] In all of the above-described embodiments, the x-ray emissionelement of the target assembly is adapted to be adjacent to or withinthe region to be irradiated. The proximity of the emission element tothe targeted region, e.g. the tumor, eliminates the need for the highvoltages of presently used machines, to achieve satisfactory x-raypenetration through the body wall to the tumor site. The low voltagealso concentrates the radiation in the targeted tumor, and limits thedamage to surrounding tissue and surface skin at the point ofpenetration. For example, the delivery of 4000 rads, as is requiredafter a mastectomy, with a 40 kV, 20 uA electron beam, may requireapproximately 1 to 3 hours of radiation. However, since the x-ray sourceis, in this preferred embodiment, insertable proximate to, or into, theregion-to-be-irradiated risk of incidental radiation exposure to otherparts of the patient's body is significantly reduced.

[0038] Further, specificity in treating tumors may be achieved bytailoring the target and shield geometry and material at the emissionsite, for example as disclosed in U.S. Pat. No. PHLL-111, assigned tothe assignee of the present invention. This tailoring facilitates thecontrol of energy and the spatial profile of the x-ray emission toensure more homogenous distribution of the radiation throughout thetargeted tumor.

[0039]FIG. 4 is a schematic representation of the x-ray source apparatus10 shown in FIG. 1A. In that preferred configuration, the housing 12 isdivided into a first portion 12′ and a second portion 12″. Enclosedwithin the first housing portion 12′ is a rechargeable battery 12B, arecharge network 12D for the battery 12B, which is adapted for use withan external charger 50, and a telemetry network 12E, adapted to beresponsive to an external telemetry device 52 to function in the mannerdescribed below. That portion 12′ is coupled by cables to the secondhousing portion 12″. The second housing portion 12″ includes the highvoltage power supply 12A, controller 12C and the probe 14, as well asthe electron beam generating element 22. In one embodiment, the electronbeam generator includes a thermionic emitter 22 driven by the powersupply 12A. In operation, power supply 12A heats the thermionic emitter22, which in turn generates electrons which are then accelerated towardthe anode 24. The anode 24 attracts the electrons, but passes themthrough its central aperture toward the target assembly 26. Thecontroller 12C controls the power supply 12A to dynamically adjust thecathode voltage, the electron beam current, and temporal parameters, orto provide pre-selected voltage, beam current, and temporal parameters.

[0040] Also illustrated, is an alternative electron beam generator whichincludes a photoemitter 22 irradiated by a light source 56, such as adiode laser or LED, powered by a driver 55. The light is focused on thephotoemitter 22 by a focusing lens 58.

[0041] In the illustrated embodiment, external telemetry device 52 andtelemetry network 12E cooperate to permit external control (dynamic orpredetermined) control over the power supply 12A and temporalparameters. In embodiments when the housing 12″ is not implanted, butwhere only probe 14 extends into a patient's body, the controller 12Cmay directly be used to control operation and in that case there is noneed for network 12E.

[0042]FIGS. 5 and 6 show a diagrammatic view of radiation treatmentapparatus 200 including a flexible probe 214. The apparatus 200 includesa high voltage source 218, a laser (or other optical) source 220, aprobe assembly 214, and a radiation source assembly 226. According toone aspect of the invention, the apparatus 200 provides the requiredflexibility, without using strong magnetic fields, by locating electronsource components 222, 223 and accelerator 224 near the target 228 inthe distal end of the probe 214. The probe assembly 214 couples both thelaser source 220 and the high voltage feed 218 to the radiation sourceassembly 226. Preferably, the probe assembly includes flexible fiberoptical cable 202 enclosed in a small-diameter flexible metallic tube204.

[0043] The radiation source assembly 226, which can be for example 1 to2 cm in length, extends from the end of the probe assembly 214 andincludes a shell which encloses the target 228. According to oneembodiment, the radiation source assembly 226 is rigid in nature andgenerally cylindrical in shape. In this embodiment the cylindrical shellenclosing the radiations source assembly 226 can be considered toprovide a housing for the electron beam source as well as a tubularprobe extending from the housing along the electron beam path. The innersurface 226A of the assembly 226 is lined with an electrical insulator,while the external surface of the assembly 226 is electricallyconductive. According to a preferred embodiment, the radiation sourceassembly is hermetically sealed to the end of the probe assembly 214,and evacuated. According to another embodiment, the entire probeassembly 214 is evacuated.

[0044] The terminal end 202A of the fiber optical cable 202 ispreferably coated, over at least part of its area, with asemitransparent photoemissive substance such as, Ag-O-Cs, thus forming aphotocathode 222. A high voltage conductor 208, embedded in the fiberoptical cable 202, conducts electrons to the cathode 222 (if necessary),the electron multiplier 223 and the accelerator 224 from the highvoltage source 218. Similarly, the flexible tube 204 couples a groundreturn from the target 228 to the high voltage source 218, therebyestablishing a high voltage field between the cathode 216 and the target228. The fiber optical cable 202 acts as an insulating dielectricbetween the high voltage conductor 208 and the grounded flexible tube204.

[0045] In order to eliminate scattering of the light in the fiber opticcable 202 by the high voltage wire 208, the fiber optic cable 202 canhave an annular configuration. The light from the laser 220 travels downthe annular core of the fiber optic cable 202. Cladding can be providedon each side of the core having an index of refraction so as to reflectthe light beam incident on the interface back into the core. Thegrounded flexible metal tube 204 can surround the outer cladding.

[0046] As in previously described embodiments, the target 228 can be forexample, beryllium, (Be), coated on one side with a thin film or layer228A of a higher impedance element, such as tungsten (W) or gold (Au).

[0047] In operation, the small semiconductor laser 220 shining down thefiber optical cable 202 activates the transmissive photocathode 222which generates free electrons 216. The high voltage field between thecathode 222 and target 228 accelerates these electrons, thereby forcingthem to strike the surface 228A of target 228 and produce x-rays. Inorder to generate, for example, 20 uA of current from an Ag-O-Csphotocathode 222 with a laser 220 emitting light at a wavelength of 0.8m, the 0.4% quantum efficiency of this photocathode 222 for thiswavelength requires that the laser 220 emits 7.5 mW optical power. Suchdiode lasers are readily commercially available. According to theinvention, the photoemissive surface which forms cathode 222 can, infact, be quite small. For example, for a current density at the cathode222 of 1 A/cm², the photoemitter's diameter need only be approximately50 μm.

[0048] One difficult fabrication aspect of this invention is thefabrication of the photocathode 222, which for practical substances,with reasonable quantum efficiencies above 10⁻³, should be performed ina vacuum. This procedure can be carried out with the fiber optical cable202 positioned in a bell jar, where for example, an Ag-O-Cs photosurfaceis fabricated in the conventional manner. Subsequently, without exposureto air, the optical cable 202 can be inserted into the tube 204. The end202B can be vacuum sealed to the flexible tube 204.

[0049] In the above embodiments, the probe 14 or 214, along with itsassociated target 26, or 228, can be coated with a biocompatible outerlayer, such as titanium nitride on a sublayer of nickel. For additionalbiocompatibility, a sheath of, for example, polyurethane can be fittedover the probe, such as that illustrated in FIG. 3.

[0050] The invention may be embodied in other specific forms withoutdeparting from the spirit or essential characteristics thereof. Thepresent embodiments are therefore to be considered in respects asillustrative and not restrictive, the scope of the invention beingindicated by the appended claims rather than by the foregoingdescription, and all changes which come within the meaning and range ofthe equivalency of the claims are therefore intended to be embracedtherein.

What is claimed is:
 1. A miniature radiation source comprising: aflexible catheter extending along a probe axis between a proximal endand a distal end of the catheter; a radiation source having asubstantially rigid housing defining a substantially evacuated interiorregion extending along a beam axis between an electron source at aninput end of the housing and a radiation transmissive window at anoutput end of the housing, the housing having a target responsive toincident accelerated free electrons to emit radiation whereby theradiation emitted therefrom is directed through the radiationtransmissive window, the housing having the input end affixed to thedistal end of the catheter; means responsive to a signal at the proximalend of the catheter for selectively activating the electron source toemit free electrons in the interior region a channel electron multiplieradapted for receiving electrons from the electron source, at an inputend, and adapted for producing free electrons at an output end; and anelectron accelerator adapted for establishing a potential difference inthe interior region of the radiation source whereby the free electronsare accelerated toward the target.
 2. A miniature radiation sourceaccording to claim 1 , wherein the electron accelerator includes asurface disposed about the beam axis between the electron source and thetarget and being characterized as having an increasing distance from thebeam axis as a function of distance from the electron source, thesurface bearing a semiconductor coating.
 3. A miniature radiation sourceaccording to claim 2 , wherein the surface is established by amonolithic ceramic element.
 4. A miniature radiation source according toclaims 2, wherein the surface is a surface of revolution.
 5. A miniatureradiation source according to claims 2, wherein the function is linear.6. A miniature radiation source according to claim 1 , wherein theelectron source is selected from the group including a photocathodeilluminated by laser energy, a field emitter, and a thermionic emitter.7. A miniature radiation source according to claim 1 , wherein thetarget is at ground potential.
 8. A therapeutic radiation source adaptedfor coupling to a catheter, comprising: a radiation source having asubstantially rigid housing defining a substantially evacuated interiorregion extending along a beam axis between an electron source at aninput end of the housing and a radiation transmissive window at anoutput end of the housing, the radiation source having a targetresponsive to incident accelerated free electrons to emit radiationdisposed along the beam axis whereby the radiation emitted therefrom isdirected through the radiation transmissive window, the electron sourcegenerating the electrons in response to a signal communicated throughthe catheter, the window being at ground potential; and a ceramicmonolithic accelerator for accelerating the electrons along the beamaxis, the accelerator having a hollow interior for passing electronstherethrough, the interior diverging away from electron source, theinterior being coated with a semiconductor coating to provide a smoothvoltage gradient along the beam axis and to reduce secondary emissionsfrom the accelerator.
 9. A therapeutic radiation source according toclaim 8 , further comprising means connecting the coating to a controlvoltage in the catheter for modifying electron acceleration through theaccelerator.
 10. A therapeutic radiation source according to claim 8 ,further comprising a channel electron multiplier adjacent to the sourcefor multiplying electrons for acceleration within the accelerator.
 11. Atherapeutic radiation source according to claim 8 , wherein the interiorcomprises a single cylindrical wall.
 12. A therapeutic radiation sourceaccording to claim 8 , wherein the electron source is selected from thegroup consisting of a photocathode illuminated by laser energy, a fieldemitter, and a thermionic emitter.
 13. A therapeutic radiation sourceaccording to claim 8 , wherein the source comprises a channel electronmultiplier.
 14. A radiation source for attachment to a therapeuticprobe, the radiation source having an elongated rigid housing defining asubstantially evacuated interior region extending along a beam axis,said radiation source comprising: an electron source for producing freeelectrons; a ceramic monolithic accelerator for accelerating the freeelectrons toward a radiation target, the accelerator having a hollowinterior for passing electrons therethrough, the interior diverging awayfrom electrons of increasing acceleration, the interior being coatedwith a semiconductor coating to provide a substantially smooth voltagegradient between high voltage and ground potential and to reducesecondary emissions from the accelerator, the radiation targetconverting accelerated electrons into radiation; an radiation window,coupled to ground potential, for passing the radiation through thehousing.
 15. A radiation source according to claim 14 where in saidelectron source comprises an electron generator and an electronmultiplier, said electron multiplier being adapted for producing aquantity of free electrons as a function of a quantity of electronsproduced by said electron generator.